With the conventional still-picture X-ray technology currently in use, the so-called film-based technique is the most prevalent. In this technique, the patient is exposed to X-rays and the X-rays that pass through the body are then exposed onto a sheet of film. The film has the function of displaying and recording information, and is widely used throughout the world due to its capacity to be enlarged, its high degree of gradation, its light weight and ease of handling. On the other hand, the technique suffers from several disadvantages, including a complicated process of developing the image, the problem of long-term storage, and the time and effort involved in manual search and retrieval of the physical images.
By contrast, moving image photographic systems rely mainly on image intensifier (I.I). Since I.I uses the photo-electron multiplier effect inside the device, it generally has good sensitivity and has the additional advantage of exposing the patient to lower levels of radiation. The I.I not only provides the physician with a see-through image of the patient but also, due to the conversion of the CCD analog output to digital output (a process referred to here as digitization), makes possible the computerized recording, display and storage of such data.
However, because medical diagnosis requires a high degree of gradation, even with I.I, film is often used for still picture imaging. In addition, such systems suffer from the following disadvantages: peripheral image distortion due to the characteristics of the optical system, low contrast and large equipment size.
Recently, with a growing need to digitize X-ray images inside the hospital itself, in place of film, X-ray imaging devices that use an X-ray sensor with solid-state image sensing elements arrayed two-dimensionally to convert the X-ray image into electrical signals have begun to be used. Since the X-ray image can then be replaced with digital information, image information can be sent instantaneously to distant locations, with the advantage of being able to provide state-of-the-art, high-quality diagnostics even to remote areas. Moreover, if no film is used the space previously required for its storage can be turned to other, more productive uses. If in the future it becomes possible to introduce more advanced and sophisticated image processing techniques, it is possible that diagnostics may to some extent be computerized and therefore automated, without the intervention of a radiologist.
Moreover, in recent years, with the use of amorphous thin-film semiconductors in solid-state image sensing elements, X-ray imaging devices capable of taking still pictures have been developed. Using amorphous silicon thin-film semiconductor production technology, photos exceeding 40 cm a side and capable of completely imaging the human torso have been commercialized. Since the production process itself is relatively simple, it is expected that inexpensive detectors based on this technique will become available in the not-so-distant future. In addition, since amorphous silicon can be produced in thin glass sheets having a thickness of 1 mm or less, the detector itself can be made very thin and compact, for greater ease of handling.
More specifically, FIG. 5 shows the internal structure of a read-out circuit of an X-ray imaging device using an amorphous silicon thin film semiconductor for the solid-state image sensing elements. In FIG. 5, RES1–RES3 are reset switches that reset the signal lines M1–M3, A1–A3 are amplifiers that amplify the signals of M1–M3, CL1–CL3 are sample-hold capacitors that temporarily store the signals amplified by amplifiers A1–A3, Sn1–Sn3 are sample-hold switches, B1–B3 are buffer amplifiers, Sr1–Sr3 are switches for the serial conversion of parallel signals, reference numeral 103 denotes a shift resister for applying pulses to the switches Sr1–Sr3 for serial conversion, and reference numeral 104 denotes a buffer amplifier for outputting the serially converted signals.
FIG. 6 is a timing chart showing the operation of an X-ray imaging apparatus having the read-out circuit shown in FIG. 5. First, as for the photoelectric conversion interval (given in the diagram as the X-ray exposure interval): In a state in which the TFT are all OFF and when the light source (X-rays) are turned ON in pulses, each of the respective photoelectric converters is struck by the light and a signal electric charge comparable to the amount of light is stored in each of the respective converter capacitors. If a fluorescent material is used to convert the X-rays into visible light, then either a light-guiding member for guiding the light made visible in proportion to the number of X-rays to the photoelectric converters may be used or the fluorescent material may be disposed near the electrodes of the photoelectric converters.
It should be noted that the signal electrical charge is held in the converter capacitor after the light source is OFF.
Next, as for the read-out interval: The read-out is accomplished at the S1-1–S3-3, one row at a time, starting with row S1-1–S1-3, then with row S2-1–S2-3, and finally with row S3-1–S3-3. First, a gate pulse is applied from SR1 to the T1-1–T1-3 (TFT) switch gate lines in order to read out the first row S1-1–S1-3. Doing so turns T1-1–T1-3 ON and the signal electrical charges that had been stored in S1-1–S1-3 is sent to the signal lines M1–M3 to which read-out capacitors CM1–CM3 (see FIG. 1) have been added, so that the signal electrical charges are sent to the read-out capacitors CM1–CM3 via the TFT. For example, read-out capacitor CM1 added to signal line M1 is the (three-) sum total of the T1-1–T1-3 gate-source interelectrode capacitance (Cgs). Amplifiers A1–A3 amplify the signal electrical charge sent to signal lines M1–M3.
The amplified signal electrical charge sent to capacitors CL1–CL3 both turns OFF and holds SMPL signal OFF. Next, by imparting a pulse from a shift resister 103 to switches Sr1, Sr2 and Sr3 (in that order) the signals held at CL1–CL3 are then output from an amplifier 104 in the order CL1, CL2 and CL3. Since analog signal outputs B1, B2 and B3 are output from the amplifier 104, the entire unit, including the shift resister 103 and the switches Sr1–Sr3, is called an analog multiplexer. Ultimately, one row's worth of photoelectric conversion signals (S1-1, S1-2, S1-3) is output in sequence by the analog multiplexer. The read-out of the second row S2-1–S2-3 and the read-out of the third row S3-1–S3-3 are carried out in the same way as the read-out of the first row described above.
If the signals at signal lines M1–M3 are sampled and held at CL1–CL3 by the first row's SMPL signal, then the signal lines M1–M3 can be reset to ground electric potential by a CRES signal and thereafter a G2 gate pulse can be applied. In other words, second-row signal electrical charges from the photoelectric converters S2-1–S2-3 can be transmitted by the SRI while at the same time the first row's signals are being serially converted by the SR2. In so doing, all the signal electrical charges of the first through third rows of photoelectric converters can be output.
Although moving image photography employing I.I has a higher sensitivity than X-ray imaging apparatuses employing amorphous silicon, as described above, I.I-based imaging suffers from the disadvantages of peripheral image distortion, low contrast and, when using photographic film, large and unwieldy equipment size.
By contrast, X-ray imaging apparatuses introducing amorphous silicon thin film semiconductors have been commercialized and provide still images of a picture quality equal or superior to that of the conventional film-based imaging method. However, there has been no commercialization of see-through moving pictures. One reason for this is that moving picture imaging exposes the patient to greater levels of radiation. Reducing the levels of radiation exposure only worsens the signal-to-noise ratio (S/N). Also, the scan speed (that is, the frame rate) of such moving picture imaging must usually be much greater (that is, much faster) than is the case with sill picture imaging, which means that the imaging apparatus's frequency band must be broadened. However, broadening the imaging apparatus's frequency band increases the so-called white noise such as shot noise and Johnson noise, thus degrading the S/N.
In FIG. 5, there are three input lines and one output line, with the parallel signals being converted to serial signals by an analog multiplexer 103 (Sr1–Sr3). Normally, amplifiers A1 and B1 are set to frequency bands at which they can amplify signals from a photoelectric converter circuit within a single line operating time interval. Broadening the band beyond what is necessary only increases the Johnson noise (that is, the thermal noise, which is electrical noise produced by thermal agitation of electrons in conductors and semiconductors), which is undesirable. On the other hand, the analog signals must be serially converted by amplifiers and multiplexer (103, Sr1–Sr3) at a later stage than that of the analog multiplexer, so the frequency band must be set broader than the frequency bands of the A1 and A3 amplifiers.
Generally, the Johnson noise (Vrms) is expressed by the following equation:Vrms=(4KTRB)1/2where K is the Boltzmann constant, T is absolute temperature, R is resistance and B is the frequency band. Portion (4KTR)1/2 unmultiplied by B1/2 is called the noise density.
An operational amplifier generates Johnson noise from all the components that compose it, such as the ON-resistance of the transistor that forms the first stage of the amplifier and the input resistance and feedback resistance not shown in the FIG. 5, such noise being proportional to the square root of the individual frequency bands.
For example, in FIG. 5, amplifier 104 serially converts three pixels, so it requires a frequency band that is at least three times greater than that of amplifier A1. This means that, as a result, when the A1 amplifier noise density and the 104 amplifier noise density are equal, the 104 amplifier noise (that is, the effective value) increases.
The chest X-rays used in hospitals, because they are so often used for chest X-rays, are said to require a photosensitive area of at least 40 cm×40 cm. In that case, the pixel pitch should be no greater than 200 μm, preferably less if possible. For example, an area measuring 40 cm×40 cm with a pitch of 200 μm would require 2,000×2,000 pixels=4,000,000 pixels. In other words, the number of line inputs in the read-out circuit in FIG. 5 would be 2,000 (pixels).
Accordingly, it is not practical to produce an arrangement calling for 2,000 input lines over 40 cm on a single read-out circuit, so ordinarily a design is used in which the load is apportioned among a plurality of read-out circuits. For example, if divided among ten read-out circuits, the number of inputs is 200 per 1 output, which is a more realistic range. In this case, the 104 amplifier requires a frequency band that is 200 times that of the A1 amplifier. Assuming the 104 amplifier and the A1 amplifier have the same noise density, the noise effective value will work out to be 2001/2≈14 times for the 104 amplifier.